Traditional optical methods of functional human brain imaging, such as diffuse optical tomography, may be used to generate images of optical properties in the brain, which may provide information about localized neural activity, e.g., via neural activity dependent hemodynamic properties such as blood flow and/or hemoglobin oxygenation state, or other neural activity-dependent optical parameters impacting light absorption, scattering or other properties. However, because diffuse optical tomography methods typically rely on light that is randomly scattered inside the brain, the spatial resolution can be relatively limited, especially at deeper depths. The light that emerges from the brain tissue and impinges on a sensor is largely composed of multiply scattered photons that have taken a wide range of highly tortuous paths in the manner of a diffusive random walk.
Optical coherence tomography (OCT) is one example of an optical technique that can be used to image at depth inside tissue. In an OCT system, light from a low-coherence source is split into two paths along two different arms of an interferometer: a reference arm and a sample arm. Various optical components may be provided along each arm to define and/or adjust specific beam parameters, such as the shape, focal depth, and/or intensity distribution of the light. In the reference arm, the light is back-reflected by a mirror and it returns into the interference system, propagating along the same path it came from but in the opposite direction. In the sample arm, the light is backscattered through the sample medium back into the interference system. The returning light from both arms recombine or merge at a coupler to generate an interference pattern. By tuning the position of the mirror in the reference arm, the optical distance for the light propagating through the reference arm can be adjusted and the interference pattern is formed only with light that traveled, to within approximately the coherence length of the light source, the same optical distance in the sample. The dependence on optical path length of the intensity of light backscattered from beneath a sample surface can be measured based on the interference patterns resulting from varying the path length of the reference arm.
OCT has typically been applied in a microscopic mode at limited imaging depth, and in such a mode, the spatial resolution in both the axial direction (along the Z-axis) and the lateral direction (across the XY plane) may range between about 1 μm to about 10 μm, but the penetration depth (i.e., along the Z-axis) of such conventional OCT is typically only about 1-2 mm. In certain embodiments of OCT (Giacomelli M G, Wax A. “Imaging beyond the ballistic limit in coherence imaging using multiply scattered light”. Optics Express. 2011 Feb. 28; 19(5):4268-79), it has been possible to image at nearly 1 cm depth with nearly 1 mm resolution. The typical thickness of a human skull, however, is from about 4 mm to about 10 mm and including the thickness of skin and any intervening dura layer, pia layer, and cerebral spinal fluid between the skull and the brain, and furthermore in order to image at depth inside the brain, an optical modality must have an imaging depth of at least 10 mm. Moreover, to operate in a tomographic mode, with source and detector located at a distance from one another on the scalp, analogous to diffuse optical tomography, an optical modality must be able to operate at several centimeters of path length between light source and light detector.
Accordingly, improvements to the penetration depth and spatial resolution of optical coherence tomography methods for imaging brain activity are desirable.